Measuring systems for biochemical analysis are important components of clinically relevant analytical methods. This primarily concerns the measurement of analytes which can be directly or indirectly determined with the aid of enzymes. Biosensors, i.e., measuring systems equipped with biological components, which allow a repeated measurement of the analyte either continuously or discontinuously and which can be used ex vivo as well as in vivo have proven to be particularly suitable for the measurement of analytes. Ex vivo biosensors are typically used in flow-through cells whereas in vivo biosensors are preferably implanted into subcutaneous fat tissue. In this connection one distinguishes between transcutaneous implants which are only introduced into the tissue for a short period and are in direct contact with a measuring device located on the skin, and full implants which are inserted surgically into the tissue together with a measuring device.
Electrochemical biosensors allow the measurement of an analyte by means of two or more electrodes where at least one of the electrodes represents the working electrode on which the analyte to be determined is converted. Electrochemical biosensors which comprise an enzyme as a biological component contain the enzyme in or on the working electrode in which case for example the analyte can serve as a substrate for the enzyme and can be physicochemically altered (e.g. oxidized) by this enzyme. A redox mediator transfers the electrons released during the conversion of the analyte onto the conductive components of the working electrode, and the electrical measuring signal generated by the flow of electrons correlates with the concentration of the measured analyte.
Naturally occurring as well as synthetic redox pairs come into consideration as redox mediators. Synthetic redox mediators such as for example those described in the publication by Feldman et al. Diabetes Technology & Therapeutics 5 (2003), 769-779 are less suitable for in vivo applications. This is due to the fact that a synthetic redox mediator can theoretically always produce an immune response by the body when the biosensor is introduced into the body. However, at least the toxicity of these substances must be considered and, if necessary, checked because redox mediators must always be able to freely diffuse through the electrode structure by which means they can also escape from the electrode and pass over into the surrounding organism. This point is not relevant for ex vivo applications provided it is ensured that it does not enter the body due to a potential return flow of the analyte.
Consequently, electrochemical sensors which use naturally occurring redox mediators are particularly suitable for in vivo applications. The redox pair oxygen/hydrogen peroxide proves to be particularly advantageous in this connection because the initial component (oxygen) is always present. The hydrogen peroxide generated in the enzymatic conversion of an analyte by means of an oxidase in the presence of oxygen is reoxidized on the working electrode of the electrochemical biosensor whereupon an electrical signal is generated by the release of electrons and the redox mediator is converted back into its oxidized form. The kinetics of this enzymatic reaction follows a so-called ping-pong mechanism. Leskovac et al., The International Journal of Biochemistry and Cell Biology 37 (2005), 731-750.
A significant problem when measuring analytes with the aid of enzymes which require oxygen as a co-substrate is, however, that temporary reductions of the oxygen concentration compared to the initial situation can occur in tissues which can affect the function of conventional in-vivo biosensors. FIG. 1 shows the kinetics of the enzymatic oxidation of glucose to glucono-δ-lactone by means of glucose oxidase at various oxygen concentrations. The graph shows that in general the amount of analyte converted at a given oxygen concentration is reduced as the glucose concentration increases and, thus, the curve is in the non-linear range in the physiologically relevant range despite the high binding constant of glucose oxidase for glucose (about 250 mM).
Furthermore, FIG. 1 shows that at higher concentrations of the analyte, an approximately linear curve is not obtained until an oxygen concentration of about 1 mM. The in vivo concentration of dissolved oxygen in aqueous systems and in particular in the interstitial fluid of subcutaneous fat tissue is, however, considerably lower. Whereas water has an approximate oxygen concentration of 0.21 mM at 37° C., the expected oxygen concentration in subcutaneous fat tissue is only 0.1 mM or even less which is why the curves are in each case curved at physiological glucose concentrations. This deviation from a linear course results in undesired transient function characteristics in in vivo biosensors.
Thus, the limited availability of oxygen in the tissue is a limiting factor for the linearity of the function curve of the electrochemical sensor in numerous enzymatic biosensors which require oxygen as a co-substrate. The linearity of the function curve can in principle be improved by using working electrodes with a cover membrane which inhibits the diffusion of the analyte more strongly than the diffusion of the co-substrate. FIG. 3 shows among others the function curve of an enzymatic biosensor with a cover membrane consisting of polyurethane which promotes the diffusion of oxygen more strongly than that of glucose (measured values indicated by squares). This shows that the measuring signal of the sensor can be kept approximately linear up to a glucose concentration of about 10 mM by using a suitable cover membrane. The curve becomes increasingly curved at higher concentrations.
However, the use of cover membranes in electrochemical sensors is associated with certain problems. Thus, electrochemical sensors which are used to determine different analytes must usually also contain different cover membranes in order to provide a different diffusion of the substrate and co-substrate. At the same time it must be ensured that the cover membranes are highly biocompatible for in vivo applications which involves considerable technical requirements and ultimately leads to increased production costs.
In order to lower the polarization voltage of the working electrode of an electrochemical biosensor relative to a reference electrode and thus to reduce the affect of interfering substances on the measuring signal of the working electrode, some electrochemical biosensors additionally utilize an electrocatalyst which promotes the transfer of electrons from the redox mediator onto the conductive components of the working electrode. An example of such an electrocatalyst is cobalt phthalocyanine which catalyses the oxidation of hydrogen peroxide to oxygen. Crouch et al., Biosensors and Bioelectronics 21 (2005), 712-718. In this process the cobalt (II) cation of the cobalt phthalocyanine complex is firstly reduced by hydrogen peroxide to cobalt (I) before it is converted back into its original divalent state with release of an electron at the anode.
Another example of an electrocatalyst known from the literature is manganese dioxide in the form of pyrolusite. Cui et al., Nanomedicine: Nanotechnology, Biology and Medicine 1 (2005), 130-135;.Luo et al., Biosensors and Bioelectronics 19 (2004), 1295-1300. Although the mechanism of the catalytic oxidation of hydrogen peroxide on manganese dioxide is not understood in detail, the potential of a working electrode with manganese dioxide as the electrocatalyst is reduced by several 100 mV compared to a working electrode without manganese dioxide. Consequently, the effect of interfering substances such as ascorbate or urea on the measuring signal is considerably reduced.
Another reason for using electrocatalysts is the damage caused to enzymes by excess hydrogen peroxide. If this substance is not sufficiently rapidly decomposed at the working electrode, a denaturation of the enzyme may occur. In order to counter this problem it was proposed in the literature to synthesize enzymes that are resistant to hydrogen peroxide, for example by mutation as shown in US 2004/0137547 A1. However, it is extremely difficult to make such modifications to an enzyme without having an adverse effect on other properties of the enzyme such as for example its enzymatic specificity. Hence, the use of electrocatalysts for conversions in which hydrogen peroxide is generated appears to be considerably superior to the above method because electrocatalysts considerably increase the efficiency of the oxidation of hydrogen peroxide and in this manner prevent excess peroxide from occurring in the electrode matrix or in its environment.
An additional problem which is associated with the formation of hydrogen peroxide in an enzymatic determination of an analyte is that hydrogen peroxide can act as an inhibitor of the analyte or of the co-substrate oxygen. This competitive inhibition depends on the concentration of hydrogen peroxide and limits the conversion of the analyte. The use of an electrocatalyst which promotes the reoxidation of hydrogen peroxide to oxygen consequently also has a positive effect with regard to the conversion of the analyte.
Various factors have to be taken into account when designing electrochemical biosensors. Thus, the biosensors have to have a sufficient amount of enzyme in the working electrode in order to prevent an enzyme limitation of the measurement. Abel et al., Journal of Molecular Catalysis B: Enzymatic 7 (1999), 93-100. Furthermore, the enzyme molecules should be located in the structure of the working electrode over the complete measuring period of the biosensor i.e. the enzyme should not become detached or displaced in areas of the electrode which are reached by the measuring medium. Doretti et al., Biosensors and Bioelectronics 11 (1996), 363-373. Finally, the enzyme should also be stable in the working electrode of the biosensor. Factors which result in a thermal deactivation of enzymes in electrochemical biosensors together with methods for their stabilization have been investigated many times. Sarath Babu et al., Biosensors and Bioelectronics 19 (2004), 1337-1341. Enzyme degradation after manufacture of a biosensor ultimately leads to a limited shelf-life of the sensor.
In order to take the above factors into consideration, attempts were made to stabilize the enzyme by immobilizing it in the electrode matrix of the working electrode which has led to an intensive search for suitable immobilization methods for enzymes in electrochemical biosensors. An adsorptive as well as a chemical immobilization are used in practice. However, adsorptive immobilization is disadvantageous for various reasons. On the one hand, it requires that the working electrode is covered by a membrane that is impermeable to the enzyme which increases the work required to manufacture the biosensor and makes a wide variety of demands on the membrane. On the other hand, the aforementioned displacement of enzyme molecules within the electrode cannot be prevented in the case of adsorptive immobilization which results in a change in the sensor function. U.S. Pat. No. 5,368,707 discloses biosensors which comprise working electrodes with an adsorptively bound enzyme and which are suitable for determining micromolar amounts of lead ions in liquids. In order to produce the biosensors, the surface of the working electrode consisting of a conductive material is coated with colloidal gold on the particles of which the appropriate enzyme is adsorbed which, in turn, can be covalently bound to a redox mediator
Another disadvantage of electrodes provided with a cover membrane for supporting the adsorptive immobilization of enzymes which should not be underestimated especially for in vivo applications, is the necessity to non-invasively check the integrity of the cover membrane. Since even the smallest defects in the membrane are sufficient to result in a bleeding of the enzyme from the electrode into the environment, an enormous amount of checking is necessary especially in the case of in vivo biosensors. Hence, in view of the disadvantages of an adsorptive immobilization there is thus a concrete need to immobilize enzymes in electrochemical biosensors by covalent bonds to or in the electrode matrix.
Japanese Patent No. JP 10-68651 describes sensors for detecting analytes such as glucose which comprise electrodes with a covalently bound enzyme. For this purpose, the surface of the electrodes coated with SnO2 as a conductive material is activated with a strong acid, functionalized with a coupling reagent and finally brought into contact with the enzyme.
European Patent No. EP 0 247 850 A1 discloses biosensors for the amperometric detection of an analyte. These sensors contain electrodes with immobilized enzymes which are immobilized or adsorbed onto the surface of an electrically conducting support where the support consists of a platinized porous layer of resin-bound carbon or graphite particles or contains such a layer. For this purpose, electrodes made of platinized graphite and a polymeric binding agent are firstly prepared and these are subsequently brought into contact with the enzyme. In this case, the enzyme is immobilized either by adsorption to the electrode surface or by coupling it to the polymeric binding agent using suitable reagents.
Amperometric biosensors with electrodes comprising an enzyme immobilized or adsorbed onto or into an electrically conducting, porous electrode material are also described in EP 0 603 154 A2. In order to produce the enzyme electrodes, an oxide or oxide hydrate of a transition metal of the fourth period, such as for example manganese dioxide, acting as a catalyst is worked into a paste together with graphite and a non-conducting polymeric binding agent, and the porous electrode material obtained after drying the paste is brought into contact with the enzyme in a second step. The enzyme can be immobilized on or in the porous electrode material by cross-linking using glutardialdehyde.
A major disadvantage of the electrochemical biosensors described in JP 10-68651, EP 0 247 850 A1 and EP 0 603 154 A2 is that the enzyme is first immobilized on the electrode that has been prefabricated without enzyme. As a consequence, there is the problem that the enzyme cannot be coupled to the electrode components in a controlled manner. Thus, when glutardialdehyde is used as a cross-linking reagent, the enzyme not only binds in an uncontrolled manner to any reactive components of the electrode material, but is also inter-crosslinked. Furthermore, this procedure contaminates the electrode with the reagents that are used and, hence, the electrode has to again be thoroughly cleaned especially before use in an in vivo biosensor which increases the production complexity and thus the costs.
U.S. Pat. No. 4,938,860 discloses a suitable electrode for electrochemical sensors comprising a platinum coated anode formed as a film and an enzyme layer which is bound to the anode. The enzyme layer is bound to the platinized anode preferably by using an aminosilane and a suitable cross-linking agent such as for example glutardialdehyde. However, a disadvantage of the electrode described in U.S. Pat. No. 4,938,860 is that due to the construction of the anode as a film only a small surface is provided for the enzymatic conversion of the analyte and platinum is a relatively expensive material to use as a catalyst.